bone regeneration

Accepted Manuscript Structure, physical properties, biocompatibility and in vitro/vivo degradation behavior of anti-infective polycaprolactone-based electrospun membranes for guided tissue/ bone regeneration Rui Shi, Jiajia Xue, Min He, Dafu Chen, Liqun Zhang, Wei Tian PII:

S0141-3910(14)00277-8

DOI:

10.1016/j.polymdegradstab.2014.07.017

Reference:

PDST 7406

To appear in:

Polymer Degradation and Stability

Received Date: 19 May 2014 Revised Date:

14 July 2014

Accepted Date: 18 July 2014

Please cite this article as: Shi R, Xue J, He M, Chen D, Zhang L, , Tian W, Structure, physical properties, biocompatibility and in vitro/vivo degradation behavior of anti-infective polycaprolactonebased electrospun membranes for guided tissue/bone regeneration, Polymer Degradation and Stability (2014), doi: 10.1016/j.polymdegradstab.2014.07.017. This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

*Revised manuscript with no changes marked Click here to view linked References

ACCEPTED MANUSCRIPT Structure, physical properties, biocompatibility and in vitro/vivo degradation behavior of anti-infective polycaprolactone-based electrospun membranes for guided tissue/bone regeneration

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Rui Shia, Jiajia Xueb, Min Heb, Dafu Chena, Liqun Zhangb※, and Wei Tiana,c※

a

Laboratory of Bone Tissue Engineering of Beijing Research Institute of Traumatology and Orthopaedics, Beijing

Beijing Laboratory of Biomedical Materials, Beijing University of Chemical Technology, Beijing 100029, PR

China c

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b

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100035, China

Department of Spine Surgery of Beijing JiShuiTan Hospital, the Fourth Clinical Medical College of Peking

*Corresponding author:

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University, Beijing 100035, China

E-Mail: [email protected] (Liqun Zhang); [email protected] (Wei Tian)

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Tel.: +86-010-64421186 (Liqun Zhang.); +86-010-58516385 (Wei Tian);

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Fax: +86-010-64433964 (Liqun Zhang); +86-010-58516538 (Wei Tian)

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ACCEPTED MANUSCRIPT Abstract: Nanofiber membranes composed of polycaprolactone (PCL), PCL/metronidazole (MNA), PCL/gelatin/MNA, and PCL/gelatin/MNA/acetic acid (HAC), named P0, P30, PG30, and PGH30,

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respectively, were fabricated by electrospinning for application in guided tissue/bone regeneration (GTR/GBR) therapies. The architectural features, mechanical properties, hydrophilicity,

drug-encapsulation efficiency, drug-release pattern, antimicrobial properties, cell barrier functions, in

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vitro/vivo degradability and biocompatibility were investigated. All membranes were found to have

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high tensile strength, which is required for GTR applications. Strong interactions among PCL, gelatin, and MNA resulted in high drug loading efficiency, which was further improved by the incorporation of gelatin and HAC. MNA incorporation gave the membranes good antimicrobial property, while reducing host versus graft reaction, improving the hydrophilicity and accelerating the degradation.

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Gelatin incorporation considerably improved cytocompatibility, while accelerating the degradation dramatically. Very low quantities （0.1% v/v with respect to polymer solution） of HAC effectively prevented the phase separation of PCL and gelatin, resulting in homogeneous nanofiber, which

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facilitates stable physical properties. The drug-release profiles of all drug-loading membranes were

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consistent with the inflammation cycle characteristics. High drug loading and trace amounts of HAC did not cause any adverse reactions, as evidenced by subcutaneous implantation. Both P0 and P30 maintained their cell barrier function in vivo for as long as 24 weeks; PGH30, for 8 weeks; and PG30, for less than 8 weeks. These findings enabled a comprehensive understanding of the influence of different compositions on the structure and performance of the membranes, thereby supporting the design of membranes with superior overall performance for GTR/GBR application. Key words: polycaprolactone; metronidazole; gelatin; electrospinning; guided tissue regeneration 2

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1. Introduction: Guided tissue/bone regeneration (GTR/GBR) techniques have been successfully applied for treating periodontal lesions and have provided the opportunity to form new bone [1]. This technique

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utilizes membranes as mechanical barriers to create a space around the defects, permitting bone regeneration to occur in the absence of competition for space by the surrounding connective tissues. Membranes used in GTR/GBR therapy must be biocompatible, have the proper degradation profile and

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adequate mechanical and physical properties, and sufficient sustained strength [2].

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Electrospun nanofibrous scaffolds/membranes more closely mimic the scales and morphologies of extracellular matrix (ECM) proteins (fiber diameters range from 50–500 nm). The inherently high surface to volume ratio of electrospun scaffolds can enhance cell attachment, drug loading, and sustained and controlled local drug delivery [3，4]. The average pore size of the electrospun nanofiber

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membrane is approximately 4–6 µm. Eukaryotic cell diameters range from 10–100 µm. It is difficult for gingival fibroblasts and periodontal cells to pass through an electrospinning membrane. These structural characteristics make the electrospun nanofibrous membranes very suitable for use as a

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GTR/GBR membrane.

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GTR membranes can be non-biodegradable or biodegradable. Non-biodegradable membranes require a second surgical procedure for membrane removal. Biodegradable membranes allowing a single-step procedure reduce patient discomfort. Collagen is the most widely used natural biomaterial for GTR application. However, there are some limitations, such as the loss of space maintenance under physiological conditions, risks of disease transmission to humans from animal-derived collagen, and high cost [5]. To overcome these limitations, polymeric GTR membranes were developed. Poly (lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymers have been widely considered for GTR 3

ACCEPTED MANUSCRIPT membrane applications. However, the accumulation of acid during degradation may significantly reduce pH, resulting in chronic aseptic inflammation [6]. Compared to PLA and PGA, PCL degradation does not produce a local acidic environment. Thus,

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because of its comparatively low cost and high mechanical strength, PCL is an attractive biomedical polymer. However, only a few studies have examined PCL-based GTR membranes [7,8]. PCL

degrades very slowly. The absorption time of PCL in vivo is approximately 2–3 y [5,9], which is too

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long for application in GTR/GBR treatment. Furthermore, its poor hydrophilicity reduces cell adhesion.

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Therefore, PCL is blended or copolymerized with other polymers before biomedical application. Among electrospun biodegradable polymers, electrospinning of hybrids, particularly natural and synthetic polymer blends, has received increasing attention [10,11]. Gelatin, the hydrolysis product of collagen, can be completely absorbed quickly. It is also biocompatible, biodegradable, and a major

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component of the native ECM. Gelatin is non-immunogenic and retains informational signals such as the arginine–glycine–aspartic acid sequence, which promotes cell adhesion, differentiation, and proliferation [12]. Thus, gelatin was combined with PCL to form a composite GTR membrane.

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However, poorly blended polymeric fibers are generated with porous or phase-segregated internal

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structures as a result of the weak molecular interactions between natural and polymeric polymers. This is particularly true when the natural biopolymer component reaches a threshold percent. Intense phase separation was observed in electrospun collagen/PCL hybrids when the PCL component was increased to 30% [13], and severe phase segregation occurred when the components were blended in equal parts for the cases of electrospun gelatin/PCL [14] and electrospun collagen/PLCL [15]. Therefore, controlling the formation of homogeneous internal structures of electrospun nanofibers by modulating the miscibility of solutions containing both natural and synthetic components is critical. Incorporating a 4

ACCEPTED MANUSCRIPT tiny amount of acetic acid may allow the originally turbid solution to become immediately clear and to be single-phase stable for more than 1 week [16]. Some structural and physical properties of the PCL/gelatin/HAC composites have been studied. However, biosecurity, degradation, and foreign body

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reaction following incorporation of acetic acid have not been reported. In addition, PCL/gelatin/HAC composites used to prepare the drug loading GTR/GBR membrane have not been determined.

The presence of periodontopathogens such as Porphyromonas gingivalis and Prevotella

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intermedia may negatively affect the success of periodontal regeneration. Therefore, it is very

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important to control and/or reduce bacterial contamination of the periodontal defect to enhance periodontal regeneration [17]. Metronidazole (MNA) has been used to treat infections for more than 45 y and is still successfully used for treating anaerobic bacterial infections. Although systemic administration of antibiotics is useful, high oral doses are necessary to achieve effective concentrations

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in the gingival fluid. However, long-term use may lead to the development of resistant bacterial strains. These limitations have led researchers to examine localized delivery of antibiotics directly at the diseased site [18]. Electrospun nanofiber membranes are used to achieve different controlled drug

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release profiles in drug delivery applications [19]. MNA has been integrated into PLA and PLGA

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nanofibers for local periodontitis treatment. This drug delivery system showed sustained drug release and significantly decreased bacterial viability [20]. In addition, significant improvement in periodontal regeneration following GTR was observed in dogs [21]. Although a derivative of MNA has been mixed with PCL to fabricate nanofibrous scaffolds, but PCL-based matrix loading of MNA for GTR/GBR application has not been reported. The effects of the interaction between the drug and polymer matrix on the physical-chemical properties, drug release profile, and long-term in vivo biocompatibility and biodegradability of the drug delivery system have not been thoroughly investigated. 5

ACCEPTED MANUSCRIPT In this study, electrospinning technology was used to fabricate 4 types of PCL-based GTR membranes composed of PCL, PCL/MNA, PCL/gelatin/MNA, and PCL/gelatin/MNA/HAC. The effects of the material components on the physical and chemical structure, mechanical and thermal

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properties, barrier function, antimicrobials performance, histocompatibility, and in vitro and in vivo degradation of different membranes were investigated in detail. The results of this study would render technology and material references for preparation of GTR/GBR membranes with excellent

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comprehensive properties.

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2. Materials and methods 2.1 Preparation of electrospun PCL-based nanofiber membranes

The materials used in this study are listed in Electronic Supplementary Information (ESI) S1. Composition and preparation of the spinning solution are stated in ESI S2. Membranes with different

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compositions were fabricated according to the conventional electrospinning procedure are also described in S2.

2.2 Characterization of PCl-based nanofiber membranes

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2.2.1 Morphology

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Membrane morphology was observed using a scanning electron microscope (SEM) (S4700, Hitachi Co., Tokyo, Japan) at a voltage of 20 kV. The membranes were coated with gold and observed under the microscope. Fiber diameter and pore size were measured using the ImageJ software from SEM micrographs at random locations (n = 100). Membrane thickness was measured using a micrometer, and its apparent density and porosity was estimated using equations eq.s1 and eq.s2 (see ESI S3) [22]. 2.2.2 Molecular dynamic simulation 6

ACCEPTED MANUSCRIPT Molecular dynamics simulation was used to investigate the effect of drug content in the polymer. The Discover and Amorphous Cell modules of the Materials Studio suite were used (Accelrys, San Diego, CA, USA) [23]. All theoretical calculations were performed using the condensed-phase

2.2.3 Thermal, crystal, and chemical structure and mechanical property

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optimized molecular potentials for atomistic simulation studies (COMPASS) force field (see ESI S4).

Fourier transform infrared spectroscopy (FTIR), differential scanning calorimetry (DSC), and

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X-ray diffraction (XRD) analyses (see ESI 5) were performed to investigate the chemical and thermal

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properties of the membranes. Tensile strength (TS) and elongation at break (EB) of the membranes both in the dry and wet states were evaluated using a BOSE ElectroForce 3200 test instrument (Framingham, MA, USA) (see ESI S5). 2.2.4 Surface hydrophilicity

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Membrane hydrophobicity was determined by measuring the static water contact angle (CA). Water CAs were measured using a SL200A type Contact Angle Analyzer (Solon (Shanghai) Technology Science Co., Ltd., Shanghai, China).

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2.2.5 Drug entrapment efficiency and drug release profile

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Entrapment efficiency was determined by dissolving a known mass of membrane in 1 mL of dimethylformamide and then added to 20 mL methanol drop by drop. After centrifugation, the liquid supernatant was analyzed by high-performance liquid chromatography (HPLC) at λmax = 310 nm. The amount of MNA was determined from the calibration curve of MNA. Entrapment efficiency was calculated as: %

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ACCEPTED MANUSCRIPT The drug release profile of the membranes was detected by soaking the membrane in phosphate-buffered saline (PBS). The membrane was cut into a circle of 2-cm diameter, accurately weighed, and then rinsed in 5 mL PBS solution (pH 7.4) in a container. The containers were sealed and

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placed in a water bath shaker at 100 rpm and 37°C. At predetermined time intervals, 2 mL PBS was removed for HPLC detection and another 2 mL fresh PBS was added. The amount of drug released was determined by HPLC.

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2.2.6 In vitro biodegradation

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The membrane was cut into a circle of 2-cm diameter (an average of three for each membrane), weighed, and soaked in 12-well plates with 5 mL PBS at 37°C. At predetermined time intervals, the remaining sample was carefully removed from the well, rinsed thrice using MilliQ water, dried at 50°C until the mass remained unchanged, and then weighed. The mass remaining vs. time was plotted to

was observed by SEM.

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obtain the degradation profile of the membrane. The morphology of the nanofiber during degradation

2.3 In vitro biocompatibility, barrier function, and antibacterial activity evaluation

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2.3.1 Biocompatibility

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The cytotoxicity of the membranes to L929 fibroblast cells, human periodontal ligament fibroblasts (hPDLFs), and Rat Osteoblast-like (ROS) cells was evaluated (see ESI 6). A Cell Counting Kit-8 (CCK-8) assay was performed to test the attachment and proliferation of L929 fibroblast cells on membranes (see ESI S7).

2.3.2 Barrier function to fibroblast cells

The in vitro barrier function of the membrane in L929 cells being used as model cells was tested as follows (Fig.1): The membrane was cut into a circle 2.5 cm in diameter, sterilized, fixed on a 8

ACCEPTED MANUSCRIPT Cell-Crown, and then placed into 24-well plate without touching the bottom of the well. L929 cells were resuspended in Dulbecco’s Modified Eagle Medium supplemented with 10% (v/v) fetal bovine serum at a density of 4.0  104 cells/mL. Next, 1mL culture medium without cells was added to the

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well from the outside of the Cell-Crown, and 900 μL culture medium plus 100 μL cell resuspension was added to the sample. After incubation at 37° under a 5% CO2 atmosphere for 1, 3, 5, and 7 days, the Cell-Crown was removed and the bottom of the 24-well plate was observed under an inverted phase

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contrast microscope (IX50-S8F2, Olympus, Tokyo, Japan) to investigate whether the cells had passed

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through the membrane to the bottom of the well or the culture medium. The other side of the membrane, which had no contact with the cells, was observed by SEM to determine whether the cells had penetrated the membrane. The barrier function of the membrane after 1 month of degradation in PBS

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was also tested as described above.

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Fig. 1 The photographs and schematic diagram of barrier function characterization of electrospun nanofiber membrane

2.3.3 Antibacterial activity

Antibacterial activity of the membranes against anaerobic bacteria Fusobacterium nucleatum (ATCC 25586, Chinese General Microbiological Culture Collection Center), which is commonly found in the oral cavity during infection, was determined using the modified Kirby–Bauer method (see ESI S8) as previously described [24].

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ACCEPTED MANUSCRIPT 2.4 In vivo biocompatibility and biodegradability

Thirty healthy adult male New Zealand white rabbits (2.5–3 kg each) were used as experimental animals. The protocol was approved by the Animal Ethical Committee of Beijing Jishuitan Hospital,

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and national guidelines for the care and use of laboratory animals were applied. Prior to initiating the animal experiment, the membranes (1.5 × 1.5 cm2) were sterilized by γ-irradiation (Isotron, Ede, Nederlands). The animals were anesthetized with isoflurane, and their backs were shaved and sterilized

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with alcohol and iodine scrub. Six par vertebral incisions (2 cm each) per rabbit were made

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approximately 1 cm lateral to the vertebral column to expose the dorsal subcutis. Subcutaneous pockets were created by blunt dissection. Each individual pocket held one membrane, and the incisions were closed using surgical sutures. All the surgeries were carried out in an aseptic field using aseptic technique. After being implanted for 1, 3, 5, 8, 12, and 24 weeks, membranes were taken out and

2.6.

Statistical analysis

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observed histologically and by SEM (see ESI S9).

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The data obtained from each membrane for results from mechanical test, mass loss value, and drug loading efficiency were averaged and expressed as the mean  standard deviation. Significant

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differences among different membranes with the same composition were determined using the t-test. Differences were considered statistically significant at a p value of <0.05. 3. Results

3.1 Surface morphology PCL-based nanofiber membranes approximately 250-μm thick were obtained through 20 h of electrospinning. The surface morphology of membranes P0, P30, PG30, and PGH30, as well as their fiber diameter distribution are shown in Fig. 2. Electric conductivity (EC) of different electrospinning 10

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liquid, pore size, and porosity values of different membranes are listed in Table 1.

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Fig. 2 SEM micrographs of PCL based GTR membranes

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Micrograph images of the pure PCL nanofiber showed a randomly interconnected structure, smooth morphology, and uniform distribution of the electrospun nanofiber with no beads formed. Following addition of MNA (P30), the electrical conductivity of the spinning solution increased, while the fiber’s diameter, pore size and porosity of the membrane decreased (Table 1). For PG30, the spinning liquid was transparent when PCL/HFIP and gelatin/HFIP solutions were mixed together and electrical conductivity of the mixed solution dramatically increased. However, after 3–4 h, the mixed solution became opaque and gradually separated into different phases upon standing and formed clear 11

ACCEPTED MANUSCRIPT sediment by 8–12 h. The phase boundary of the solution is indicated in Fig.S2 (see ESI S10) PG30 as a red arrow. During gelatin incorporation, fiber diameter distribution became non-uniform. Based on the fiber diameter, the fibers can be divided into thick (1.51 ± 0.3 μm diameter) and thin (0.30 ± 0.1 μm

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diameter) groups. Phase separation led to non-uniform fiber morphology. For PGH30, a very small amount of HAC (0.1% of the total spinning liquid volume) was added to the spinning solution. The mixture immediately changed from turbid to clear, even after 24-h

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incubation under ambient conditions (see Fig.S2 in ESI S10,). The resulting fibers were smooth and

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uniform, with an average fiber diameter of 0.97 ± 0.20 μm (Table 1).

Table 1 Electrical Conductivity of the Spinning Solution as well as the Pore size and porosity of membranes Pore size （μm）

P0 P30 PG30 PGH30

6.04±0.64 3.04±0.50 3.21±0.80 4.01±0.70

Porosity (%)

Electrical Conductivity(μs/cm)

76.4 70.2 70.5 70.9

6.43±1.32 7.41±1.48 18.63±2.25 19.04±1.39

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Sample

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The mean pore size of nanofiber membranes ranged from 3–6 μm. The pore size of P0 was 6.04 ± 0.64 μm, with a porosity of 76.4%. When mixed with other components, pore sizes decreased to 3–4

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μm, while porosity decreased to approximately 70%. Several small particles on surface of the drug loading membranes were observed. The number of crystal particles on PGH30 appeared to be lower than that on PG30.

3.2 Molecular dynamics simulation of hydrogen bonds (HBs) between PCL and MNA HBs between the MNA and PCL matrix of nanofiber membranes with different MNA content were simulated by molecular dynamics simulation using the Materials (Accelrys, San Diego, CA, USA)Studio software (Accelrys, San Diego, CA, USA). The PCL chains and MNA molecules were 12

ACCEPTED MANUSCRIPT built in an amorphous cell to calculate the number, length, and angle of (Accelrys, San Diego, CA, USA)HBs, with the MNA to PCL weight to mass ratio fixed at 1, 5, 20, and 40%. As shown in Fig. S3 in ESI S11, , the oxygen and nitrogen atoms of –OH, C=N, and –NO2 on MNA can hydrogen bond

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with the hydrogen atom of –CH on PCL, and the hydrogen atom of –NO2 and –OH on MNA can form hydrogen bonds with PCL ester groups. Simulation showed that at low drug loading (5%), MNA was homogeneously dispersed in the polymer, but as drug loading increased to over 20%, MNA molecules

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generally formed hydrogen bond with each other and aggregated.

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3.3 Chemical, crystallization, thermal, and mechanical properties

Fig.3. Chemical, thermal, and mechanical properties of PCL-based nanofiber membranes: (a) FTIR spectra, (b) DSC thermograms, (c) XRD patterns, and (d) stress-strain curves of membranes in wet state

FTIR was used to investigate HB interactions among different components and to determine whether a chemical reaction had occurred. During spinning, DSC and XRD measurements were used to 13

ACCEPTED MANUSCRIPT investigate PCL and MNA crystallinity, drug distribution, and acetic acid-mediated homogeneity of the nanofiber membranes (Fig. 3). TS and EB of the membranes both in the dry and wet states were measured (Table 2).

Dry State

Samples

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Table 2. TS and EB of the different PCL-based membranes in dry and wet state Wet State

EB(%)

TS (Mpa)

EB(%)

P0

9.99±1.02

325±35

12.55±2.82

152±16

P30

6.33±0.84

99±27

11.84±1.27

52±16

PG30

4.14±0.90

19±3

6.23±1.53

13±4

PGH30

4.26±1.17

24±6

3.88±0.49

30±9

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TS (Mpa)

FTIR results showed that the characteristic peaks of MNA on the spectra of P30, PG30, and PGH30 did not differ from those of pure MNA crystals, demonstrating that the electrospinning process did not adversely affect the molecular structure of MNA, particularly the antibacterial -NO2 group.

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Acetic acid did not react with other components during electrospinning. Shifting of the original absorption bands towards the lower wave numbers indicated that HBs existed among PCL, gelatin, and

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MNA molecules, which is consistent with the results of molecular dynamics simulation. DSC and XRD results demonstrated that the melting enthalpy (∆Hm) and crystallinity of the PCL matrix decreased

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(Table.S1 in ESI S12, ) after MNA incorporation. When gelatin was added, ∆Hm and PCL matrix crystallinity further decreased. However, ∆Hm and MNA crystallinity increased, while the melting point increased. For PGH30, ∆Hm and PCL matrix crystallinity decreased and ∆Hm of MNA also decreased unlike that in the case of PG30 (see ESI S12). TS and EB of the MNA drug loading membrane (P30) was lower than the values for pure PCL (P0). With the incorporation of additional gelatin, mechanical strength further decreased. Though the miscibility between PCL and gelatin in PGH30 was better than that in PG30, the mechanical properties 14

ACCEPTED MANUSCRIPT showed no obvious changes. TS of dry and wet PGH30 were similar, while other membranes showed slightly higher TS in the wet state than in the dry state. 3.4 Surface hydrophilicity

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We measured the static water CA and observed the shape of a water droplet on the nanofiber membrane surface at 5 s (Fig.S4 in ESI S13). CA values are listed in Table 3. Water CA measurement results confirmed that both MNA and gelatin can significantly improve the hydrophilicity of the

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nanofiber membrane; the CA value of PGH30 was increased compared to PG30, suggesting a slight

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decrease in hydrophilicity.

Table 3 Water contact angle values of electrospun PCL-MNA nanofiber membranes CA(°)

P0

129.64

P30

83.99

PG30

36.69

PGH30

61.16

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Samples

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3.5 Drug encapsulation efficiency and drug release profiles

Calculated encapsulation efficiency values for P30, PG30, and PGH30 were 83.2 ± 1.1, 84.9 ± 1.5,

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and 85.7 ± 0.9, respectively. All the membranes achieved a high drug loading efficiency of above 80%, demonstrating that both the PCL and PCL/gelatin composite nanofibers are good carriers for hydrophilic drugs. Incorporation of gelatin and trace amounts of acetic acid contributed to drug loading.

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Fig.4 The cumulative drug release percentage of membranes at different soaking time in PBS.

Fig. 4 shows the release profiles of different nanofiber membranes with the same amount of MNA. All the samples released above 90% MNA within 1 week and showed a large initial burst release. Subsequently, drug release was linear from 7–14 days. The cumulative drug release of all the

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membranes reached more than 97% in 14 days. After a period of 7 days, the drug release rates slowed until accumulated drug release was nearly steady; thus, almost no variance is observed in the figure. The drug release profiles of P30 and PG30 were similar. However, the drug release rate of PGH30 was

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slower than those of both P30 and PG30 from days 5–14.

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3.6 In vitro degradation behavior

In vitro degradability of membranes was also investigated gravimetrically (Fig.5a). Morphological

changes in nanofiber membranes after 30 days of degradation were observed by SEM (Fig.S5 in ESI S14). A local enlarged image is placed in the upper right corner of each SEM image. SEM morphology revealed that the overall gross morphology of P0 after degradation for 30 days showed no significant difference. However, for P30, PG30, and PGH30, a small number of microcracks and voids were observed on the nanofiber surface. Surface roughness increased and fiber fractures were present, unlike 16

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that in the pre-degradation samples (Fig. 5b).

Fig.5 mass loss during in vitro degradation of electrospun PCL-based nanofiber membranes

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In the first 3 months, the hydrolysis rate of P0 was very slow; only 10 wt% of the material degraded, with the most mass lost after 2 months (Fig. 5a). For the other samples, the mass quickly decreased during the first 2 weeks of degradation. For P30, the calculated drug percentage was 19.2% (calculation method is in ESI S14), which was similar to the value after 2 weeks of degradation. Thus,

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the mass decrease of P30 was mainly caused by drug release during the first 2 weeks. Because of the high water solubility of gelatin, the PG30 degradation rate was much higher than that of P30. The

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calculated total weight percentage of the gelatin and MNA was 59.1% (see ESI S14). Additionally, the mass decreased by 52% after 3 months of degradation, which is similar to the total mass percentage of

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MNA and gelatin. The mass loss of PGH30 after 3 months was 59.6%, which is also similar to the total mass percentage of MNA and gelatin. These results demonstrate that drug release and gelatin dissolution were the main factors affecting degradation in the first 3 months in vitro. The mechanical properties of the nanofiber membranes after degradation for 30 days were also tested to determine whether the membrane could prevent fibroblast cell infiltration into the tissue defect site and support new tissue growth. The results are shown in Table S2 (see ESI S15). TS only showed a slight decrease after 1 month of degradation, which was maintained at approximately 2–6 17

ACCEPTED MANUSCRIPT MPa. This satisfied the requirement for the GTR/GBR membrane [25, 26]. 3.7 Cell experiments 3.7.1 Cytotoxicity

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After 72h, L929 cells incubated in the extract substrates of different electrospun membranes presented spindle, triangle, and quadrangle shapes (see Fig. S6 in ESI S16), which revealed very good cellular growth. The high MNA concentration (even in the extract substrates of membranes with a high

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MNA content of 30%) did not affect cell growth. After 72 h of incubation, cells in the PG30 leaching

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solution showed higher relative growth rates (RGRs) than those in the other samples, and even higher than those in the negative control. The RGR of PGH30 reached 93% (Fig.6). Trace amounts of acetic

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acid were not cytotoxic, but the cytotoxicity of PGH30 was grade 1 (see eq.S3 in ESI S16).

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Fig.6. RGR/% of the L929 cells cultured in the extract substrates of PCL-based electrospun membranes.

Based on these results and because of the prevalence of osteoblasts and hPDLFs around tooth defects, we tested the cytotoxicity of P30 with hPDLFs and ROS cells. The fluorescence intensity of hPDLFs and ROS cells cultured for 24 h in the extract solution are shown in Fig.S7 in ESI S17. Compared with controls, the RGRs of both the cell types were >100%, indicating that the membrane did not adversely affect the viability of hPDLFs and ROS cells. 3.7.2 Attachment and proliferation of L929 fibroblast to membranes 18

ACCEPTED MANUSCRIPT A CCK-8 assay was performed to examine the influence of MNA on the adhesion and proliferation of L929 cells on different electrospun nanofiber membranes. The RGR of the L929 cells adhered on the nano-fibrous membranes and TCP (the control) after being seeded for 4 h are shown in

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Fig. 7a. L929 cells adhered to all of the PCL-based membrane well. L929 cells attached to P0 and P30 showed no significant difference and were fewer cells than those attached on TCP; however, cells adhered on PG30 and PGH30 are nearly as same as those on TCP. Gelatin addition significantly

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improved cell attachment. Cells attached on PG30 and PGH30 showed no statistical differences,

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demonstrating that HAC contained in PGH30 did not adversely affect cell attachment.

Fig.7

RGR of L929 cells (a) adhered for 4 hours and (b)proliferated for 1, 3, 5, and 7 days on

PCL-MNA nanofiber membranes. SEM micrographs of L929 fibroblasts cultured on (c) membranes with different components for 1 day and (d) for 7days

RGRs of L929 cells proliferated for 1, 3, 5, and 7 days on the electrospun nanofiber membranes

tested at 450 nm using the CCK-8 assay are shown in Fig. 7b. The number of cells increased continuously over 7 days of culture for pure PCL (P0) and the MNA loading membrane (P30). The proliferation rate of L929 on P30 was faster than that on P0. Gelatin greatly increased L929 cell 19

ACCEPTED MANUSCRIPT proliferation, as the cell number on the surface was sufficiently high on the first day, and the RGR of L929 cells changed very little over the 7 days for PG30 and PGH30. HAC contained in PGH30 appeared to have little influence on cell vitality from 1 to 7 days when compared with PG30, but the

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cell number on PGH30 remained sufficiently high to reach an estimated confluency of approximately 90%.

To observe the effects of different components on fibroblasts, SEM studies were carried out on

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membranes on days 1, 3, 5, and 7. The cellular morphology on different membranes on the 1st and 7th

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days is shown in Fig. 7c and Fig. 7d. A larger number of interactions and contacts were present among cells on PG30 and PGH30 than on P0 and P30 on the 1st day. On the 7th day, only 50–60% L929 cell confluence was observed, and the cells were primarily round shaped cells on P0 and P30, but reached an estimated confluency of approximately 90% on PG30 and PGH30 (see ESI S18).

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3.8 Barrier function to fibroblast cells

In order to test the cell barrier function of the PCL-based membranes, we designed a novel method in vitro (See Fig. 1). SEM Micrographs of the surface of the opposite side which had no contact

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with the cells after culture for 3 days are shown in Fig. 8. For all the PCL-based membranes, even after

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seeding cells for 3 days, no L929 cells were observed on surface of this side perfectly illustrates that the membranes had good cell barrier function. The same results were obtained after 1 month of degradation.

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Fig.8 The SEM micrograph of the opposite side of the cells-seeded membranes after seeding cells on

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the electrospun PCL-MNA nanofiber membranes for 3 days.

3.9 Antibacterial function

The growth of F. nucleatum was visualized directly on plates to assess antibacterial activity (Fig.

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9a). The inhibition zone diameters of the membranes at different incubation times are displayed in Fig. 9b. Agar plates with P0 membranes were completely covered with bacteria after incubation, but agar

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plates with drug-loaded membranes showed inhibition of bacterial growth that persisted for the 30 days of the incubation. Inhibition zone diameters of gelatin-containing samples were larger than that for the

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without gelatin samples over 30 days. The inhibition zone decreased as an additional drug was released; however, even after one month, there was sufficient amount of bioactive drug released to substantially inhibit bacterial growth.

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Fig.9 Inhibition of bacterial growth on agar plates: (a) Inhibition zone surrounding the PCL-based

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membranes after incubation bacterial for 30 days in anaerobic conditions at 37℃; (b) The diameter of inhibition zone of the different PCL-based membranes to the bacterial at different incubation time.

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3.10 In vivo experiment: tissue compatibility and biodegradation evolution

During the experiment, all the rabbits remained in good health and showed no wound complications. For all the membranes, no acute or chronic inflammation or necrosis and adverse tissue reactions were identified at any time period at the implant sites. Hematoxylin and eosin staining results

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are shown in Fig. 10. Figure 11 shows the longitudinal section morphological changes of these membranes at weeks 1, 12, and 24 after implantation. For all the 4 membranes, degradation was faster

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in vivo than in vitro.

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Fig.10 Histological micrographs of PCL-based nanofiber membranes with H&E staining (length of

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scale bar =100 μm)

One week post operation, no neutrophils was observed around the 4 kinds of nanofiber

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membranes, indicating that no obvious acute inflammation occurred. Some mononuclear phagocytes as well as a small number of capillaries and granulation tissue gathered around the membranes. Fibroblasts in a particular orientation can be observed in the surrounding tissue, illustrating that the fibrosis/fibrous capsule was forming. The gross view and longitudinal section of the P0 biopsy show an integrated structure after 1 week of embedding. SEM (Fig. 11) showed that the diameter and distribution of the P0 nanofibers had not changed. Compared to P0, the membrane surface began to degrade and the tissue started to migrate inside of P30 (indicated by arrow). There were more blood 23

ACCEPTED MANUSCRIPT capillaries and granulation tissues around P30 than around P0. SEM image showed a large number of red blood cells distributed in the nanofibers. For PG30, the membrane showed a translucent appearance. Degradation occurred not only on surface, but also inside of the membrane as per the histological and

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SEM observations. Some collagen fibers were intertwined with the nanofibers, and the nanofiber diameter decreased. Because gelatin dissolves very quickly, vascular components quickly grew into the membrane, and some histocytes was transported inside of the membranes as ingrowths of blood vessels,

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leading to both surface and bulk degradation (indicated by arrow). Tissue section images reflected that

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the number of tissue cells inside of PGH30 was lower than that in PG30, demonstrating that

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degradation of PGH30 in vivo is slower than that of PG30.

Fig.11 Morphology of Longitudinal section of the PCL-based membranes after subcutaneous implantation: SEM images at 1 , 12 and 24 weeks

Three weeks after subcutaneous preparation, the number of monocytes around the membrane

greatly increased. Granulation tissue, new thin-walled blood vessels, and fibrous capsule developed, accompanied by a small number of macrophages. For P0, the internal migration phenomenon of the fibrous tissue around membrane appeared to be more obvious, and further degradation appeared on the 24

ACCEPTED MANUSCRIPT material surface (indicated by arrow) compared with that after 1 week. P30 appeared similar to P0 after 3 weeks. However, PG30 showed obvious and stratified degradation both on the surface and inside of the membrane (indicated by arrow). Compared with PG30, the degradation speed of PGH30 was

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slightly slower. The hollow and stratified phenomenon was not observed in the middle region of the PGH30, as observed on the image of PG30. Fibroblasts and phagocytes did not penetrate into the membrane, and PG30 and PGH30 retained their barrier function after 3 weeks.

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After degradation for 8 weeks, P0 was surrounded by a large number of fibroblasts, monocytes,

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and multinucleated giant cells, as well as new blood vessels and granulation tissues. Some hollow space inside of the membrane appeared (indicated by arrow), but no cells were inside P0. Membrane thickness and the appearance of P30 were similar to those of P0. The barrier function of P0 and P30 were well-maintained at 8 weeks. Degradation was also observed in the inner part of the P30

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membrane (indicated by arrow). However, few microphages were around P30, similarly to P0. P30 induced the formation of a much thinner inflammatory fibrous capsule than P0 did. For PG30, the material was replaced by a fibrous tissue. A mass of fibroblasts, monocytes, and multinucleated giant

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cells, as well as blood vessels were observed in the inner part of the membrane. Integration of the

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interface between the membrane and tissue was achieved. The blocking function of PG30 completely disappeared at 8 weeks. Compared to PG30, although a large amount of hollow space appeared inside the membrane, the cells were not present throughout the membrane. Again, the degradation rate of PGH30 was slower than that of PG30. In tissue section images at 12 weeks, fibrous capsules (indicated by arrow) were observed near the P0 and P30 became thicker, and the membranes were thinner than those at 8 weeks. The vascular tissue around the membranes increased in richness. A very large number of macrophagocytes accumulated 25

ACCEPTED MANUSCRIPT and covered the membrane surface. For PG30, only some pieces of the material (indicated by blank circles) were observed. The PGH30 was also degraded into some debris after 12 week. After 24 weeks, a large number of macrophages gathered (indicated by arrow), and only segments

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that had been covered by macrophages could be observed. Even 24 weeks after implantation, the integrity of both P0 and P30 was well maintained. Based on the cross-section images of P0 and P30, there was no infiltration of cells inside the membranes until 12 weeks. However, the fibrous tissue had

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already grown into the PG30 and PGH30 after 12 weeks. Only a small amount of nanofibers distributed

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in the fibrous tissue could be observed on the SEM image (Fig. 11). The nanofibers nearly completely disappeared. Collagen fibers replaced the membrane materials after 24 weeks, according to the SEM images. 4. Discussion

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Four types of electrospinning membranes containing PCL as the main component for GTR application were examined in this study. Incorporation of different components showed that a series of changes in the macroscopic and microscopic structure occurred, resulting in different physical

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properties and related biological effects.

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4.1 Macro- and micro-structure changes as the result of different components Generally, when the conductivity of a solution increases, a larger number of electric charges are

carried by the electrospinning jet. Thus, higher elongation forces are imposed on the jet under the electrical field. Thus, the jet path becomes longer, and more stretching of the solution is induced [27]. Moreover, bending instability can be increased during electrospinning. Both higher elongation forces and greater bending instability will result in fibers with lower diameter.

26

ACCEPTED MANUSCRIPT PCL, a nonionic synthetic polymer, when dissolved in organic solvent, does not ionize in solution. By adding MNA, the viscosity and surface tension of the PCL solution decreased, while electric conductivity increased. Thus, the mean nanofiber diameter of P30 was lower than that of P0. Generally,

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smaller fiber diameter facilitates close fiber stacking. Therefore, the average pore size of P30 decreased. For PG30, both gelatin and MNA were mixed with PCL in the spinning solution. Gelatin is a

polyelectrolyte polymer with many ionizable groups such as amino and carboxylic groups, which

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produce ions when dissolved [28]. Thus, the solution containing gelatin will have a higher charge

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density on the polymer surface, and the nanofiber diameter will be thinner. However, the fiber diameter distribution of PG30 was non-uniform, and could be divided into thick and thin groups, which occurred because of phase separation during electrospinning. PCL/gelatin solution was homogeneous immediately after mixing, but became opaque and gradually separated into different phases after

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standing. For the equal mass blend (PCL/gelatin w/w 50:50), phase separation occurred as early as during 3–4 h of standing and formed clear sediment at 8–12 h, which was in accordance with the results of previous studies [16]. We used a spinning speed of 1 mL/h and approximately 20 h was

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required to fabricate a film approximately 250-μm thick. Thus, the composition and structure of PG30

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gradually changed during electrospinning, resulting in an uneven distribution of diameter (Fig.2 PG30), which leads to unstable performance. In industrial settings involving large-scale and prolonged time for nanofiber production, this limitation must be addressed.

To overcome this issue, a very small amount of HAC was applied to demonstrate the miscibility of the gelatin phase and PCL phase. The mainly reason may be that the incorporation of PCL solution to the gelatin solution will change the pH of the original solution with respect to the gelatin’s isoelectric point [29]. Gelatin is a typical protein of polyampholyte containing both amino groups and 27

ACCEPTED MANUSCRIPT carboxyl groups, and it has the potential capability to respond to the environmental stimulus. As a solvent, HFIP exhibits a stronger acidic character, so gelatin exposes protic groups on the surface and by shields apolar aprotic groups in the interior of the protein in HFIP. However, when the gelatin/HFIP

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and PCL/HFIP solutions are mixed together, because of exposure to the highly hydrophobic PCL surrounding condition, the gelatin molecules are forced back, exposing their hydrophobic groups so as to interact with the PCL molecules, leading to a slight shift of pH value and zeta potential close to the

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IP, so the gelatin molecules intent to form compact resembling particles, and may adhere together and

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form aggregates of increased sizes, which may be large and heavy enough to settle out to the bottom of the bottle from surrounding medium, and resulting in composition and structure of the fibers changed with time. However, by introducing a tiny amount of acid (HAC) to this mixed solution, gelatin is charged positively because of the protonation of amino groups. Gelatin molecules become mutually

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exclusive, and the stretched gelatin molecular chains could penetrate into the PCL phase and entangle with PCL chains, forming a miscible and transparent solution. Therefore, the most likely molecular level interpretation of the acid-mediated phase stability and solution homogeneity achieved, so

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compositionally homogeneous hybrid nanofibers was obtained by just adding a tiny amount of HAC to

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the electrospinning solution (Fig.2, PGH30). The compositionally homogeneous of the spinning solution is significant for obtaining products with stable performance.

In addition to macro-structure differences, the micro-structure also changed with the introduction

of MNA, gelatin, and HAC, including the misibility between gelatin and PCL, interactions between the polymer matrix and drug molecules, PCL crystallinity, and drug distribution. The solubility and compatibility of drugs in the drug/polymer/solvent system were determining factors for the electrospun fiber formulation with high drug encapsulation and constant release of the 28

ACCEPTED MANUSCRIPT drugs [30]. For P30, PCL macromolecular chains can form different types of HBs with MNA, improving the compatibility and solubility between PCL and MNA. Although PCL is nonpolar and MNA is polar, these substances show good compatibility because of HB interactions. In case of MNA

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exiting, electrospinning solution conductivity increased, leading to the formation of thinner nanofibers compared to pure PCL nanofibers (P0); thus, PCL crystal formation in P30 was more difficult than that in the PCL nanofibers. Additionally, MNA acted as a plasticizer to restrain PCL crystal stack formation,

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decreasing PCL crystallization. For the PG30 composite material system, the HB type appears more

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complicated. Gelatin is composed of 18 types of amino acid proteins. Because there is no regulate repeat segment structure like synthetic polymers, it is hard to use molecular dynamics simulation to predict molecular structure and hydrogen bond types. In addition to HBs between PCL and MNA observed by molecular simulation, HBs between PCL/gelatin and MNA/gelatin also existed. Because

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the oxygen and nitrogen atom of -OH, -CN, and -NO2 on MNA can HB with the hydrogen atom of -CH, -NH, and-OH on gelatin, and the -NH and -C=O on gelatin can also form HBs with the hydrogen atom of -CH and -OH on MNA, there are more groups that can HB with MNA on gelatin chains than on

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PCL. Thus, MNA molecules generally aggregate in the gelatin phase. Aggregation induces an increase

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in MNA crystallinity according to the XRD and DSC results (Fig.3). Because the electrical conductivity of the gelatin phase is higher than that of PCL, the gelatin phase is generally distributed outside of the nanofiber, and there is more MNA aggregated in the gelatin phase than in the PCL phase; thus, MNA is more likely to migrate to the nanofiber surface during fabrication, and thus more MNA crystal particles are observed on the PG30 surface than on the other samples. In the case of PGH30, the miscibility of PCL and gelatin improved; thus, the MNA distribution in PGH30 was more homogeneous than in PG30, such that there were fewer particles on the PGH30 surface than on PG30 29

ACCEPTED MANUSCRIPT (Fig.2). Additionally, because of the high miscibility, intramolecular PCL crystallization and MNA aggregation decreased, resulting in lower PCL and MNA crystallization in PGH30 than in PG30. In conclusion, different composites will result in different interactions among various components,

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resulting in alternative connections and distributions between the drug and polymer matrix. 4.2 Physical property differences

For membrane application in GTR/GBR therapies, mechanical properties, surface hydrophilicity,

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release properties are very important for inhibiting infection.

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and biodegradability are 3 of the most important physical properties. Additionally, drug loading and

For application in tissue regeneration, a porous membrane must provide sufficient mechanical properties such as stress bearing during the surgical procedure (handling and suturing) as well as the ability to withstand in vivo stresses and avoid excessive new-tissue deformation. Additionally, the GTR

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membrane must typically be soaked in saline before the implant surgery to mimic the body-fluid environment, making mechanical properties in the wet state very important. Addition of MNA may increase spinning solution conductivity, decreasing fiber diameter and TS. This can lead to the dense

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packing of the nanofibers, decreasing the relative displacement between different nanofibers and thus

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decreasing EB. For PG30 and PGH30, the low mechanical strength of gelatin further reduced the mechanical strength compared to P30. After soaking the membrane in water, some drug was released and the PCL molecule chains were rearranged, resulting in recrystallization and increased TS. This explains why TS in the wet state is slightly higher than that in the dry state . With respect to the mechanical requirements of GTR membranes, TS values of 0.017 and 2–3 MPa are sufficient [25,26]. TS of PCL-based membranes in our study ranged from 4.1–10.0 MPa in the dry state and 3.9–12.6 MPa in the wet state (Table 2). The TS of our PCL-based membranes were 30

ACCEPTED MANUSCRIPT sufficient. The hydrophilicity of biomaterials largely influences cell adhesion and proliferation [31-33]. Because of the hydroxyl on MNA and polar imidazole ring functional groups, as well as amine and

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carboxylic functional groups on gelatin, incorporation of MNA and gelatin may improve membrane hydrophilicity (see Table 3). The hydrophilicity difference between PG30 and PGH30 were attributed

to differences in homogeneity. The incorporation of acetic acid improved nanofiber homogeneity, such

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that the gelatin content on the top side of PGH30 was lower than that on PG30 fabricated in the

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gelatin-rich phase, resulting in a higher CA value for PGH30 than for PG30. Additionally, more MNA remained in the center of the PGH30 fiber than in PG30, and therefore, the drug release rate of PGH30 was slower than that of PG30. PGH30 is more suitable for sustained-release of hydrophilic drugs than PG30 is.

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In vitro degradation results (see Fig.5) showed that the hydrophilic drug and gelatin were important for degradation. Because of the very slow degradation rate of PCL within 3 months, the mass loss was mainly attributed to the drug and gelatin. Although the TS of all the membranes decreased

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slightly, new tissue regeneration can still be supported [26]. In this study, the low degradation rate of

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PCL, which limits the clinical use of PCL, was significantly increased by the incorporation of the hydrophilic MNA and gelatin. Drug encapsulation efficiency is highly influenced by the solubility and compatibility of the drug

in an electrospun solution. Although MNA is hydrophilic and PCL is hydrophobic, strong HBs can still form on improving their compatibility, according to molecular simulation and FTIR results. Additionally, MNA is highly soluble in the electrospun solution. A relatively high entrapment efficiency of 83.2% was observed for P30. Because the interactions between MNA and gelatin are 31

ACCEPTED MANUSCRIPT stronger than those between MNA and PCL, the entrapment efficiency of gelatin-containing samples should be higher than those without gelatin. However, because the drug is more likely to aggregate on the surface of the gelatin-containing samples, the improvement in the entrapment efficiency was

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limited. The HAC-mediated membrane showed more even drug distribution than the membrane without HAC did, and this improvement is also helpful for improving drug encapsulation efficiency. It has been demonstrated that both the amount and distribution of amorphous drug and drug

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particles have a profound influence on the release kinetics [34] Higher amorphous concentrations in the

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membrane lead to higher release rates and stronger initial bursting, and distributing more drug particles close to or at the membrane surface has the same effect [35].Compared to P30, PG30 preserved higher drug content. Besides, more drugs distributed in the gelatin phase. The degradation rates of gelatin phase is quicker than PCL phase. Both of the high drug content and degradation rates resulted in faster

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drug release of PG30 than P30. Because of the MNA distribution in PGH30 is more homogeneous than that in PG30, the MNA stayed in the center part of the PGH30 fiber is more than that in PG30, so the drug release rate of PGH30 is slower than that of PG30 (see Fig.4), so PGH30 is more suitable for

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sustained release of hydrophilic drugs.

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When a material is implanted into a tissue-defect site, infection and inflammation most frequently occurs during the first week. A large initial burst of antibiotic drugs is ideal for eliminating intruding bacteria before they begin to proliferate in the first few days after material implantation. Few microorganisms can survive after an initial burst; however, continued drug release is necessary for preventing the proliferation of their further population and for inhibiting the occurrence of latent infection [36]. The drug release profile of our MNA loading nanofiber membranes adapted to the inflammation cycle characteristics. Additionally, the concentration of the drug released during the first 32

ACCEPTED MANUSCRIPT week was sufficiently high to inhibit bacterial growth, but was much lower than the systemic treatment dosage (see ESI S19). 4.3 Biological property changes

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In addition to structural and physical properties changes, different material components will also cause biological changes such as those in cellular compatibility, anti-bacterial properties, histocompatibility, and in vivo degradability.

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Cells in contact with a surface will first attach, adhere, and spread. One side of the GTR/GBR

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membrane contacts the defect, while the other side contacts the fibrous tissue. Therefore, fibroblast adhesion and proliferation is helpful for wound healing and maintaining the space for defective tissue growth. According to previous studies, hydrophobic surfaces show lower cell adhesion in the initial step of cell culture, and incorporating MNA increased the hydrophilicity of the nanofiber membranes.

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Cells then proliferated on the rough surface [37]. Drug crystals may be exposed to the fiber surface during degradation, producing a rough nanofiber surface. These factors result in the faster proliferation of L929 on P30 than on P0. Additionally, a large number of hydrophilic groups such as –NH3 and

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–COOH, as well as a number of cell-recognition domains on the gelatin chains, greatly improved cell

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attachment and proliferation. Although fewer cells proliferated on the surface of PGH30 than on PG30 within 7 days, the small amount of acetic acid slightly influenced the cytotoxicity of the membrane. However, the number of cells on the surface of PGH30 was higher than that in case of P30, and the RGR increased to 93% on the 7th day, demonstrating that PGH30 is still an excellent cell-friendly material (see Fig.7). Bacterial inhibition experiments (see Fig.9) were conducted to examine the effects of matrix components on the antibacterial properties of different PCL-based nanofiber membranes at the same 33

ACCEPTED MANUSCRIPT drug-loaded contents. Because gelatin degrades very quickly, drug present in the gelatin phase will quickly diffuse onto the plate; thus, the initial antibacterial ability of PG30 and PGH30 was better than that of P30. Because of the strong interaction between drug and gelatin, most drugs exist in the gelatin

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phase. Thus, PG30 and PGH30 released more drug than P30 did, resulting in the stronger antibacterial property.

To assess the inflammatory reaction of these nanofiber membranes and determine whether a local

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high concentration of MNA and HAC contained in PGH30 could mitigate an in vivo host response, we

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performed a subcutaneous implant study (Fig.10 and11). MNA addition improved membrane hydrophilicity, increasing the degradation speed. In addition, MNA may prevent histiocyte and monocyte aggregation, and thus incorporation of MNA may restrain host reaction to some extent. The degradation rate of membranes directly affected the maintenance of barrier function and

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tissue regeneration space. As a result of in vivo implantation, the typical foreign body response occurred, which involved the accumulation of cells, such as macrophages. free radicals, acidic products, or enzymes produced by these cells during the foreign body response; these factors may accelerate

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degradation [38], increasing the in vivo degradation rate compared to the in vitro rate. With the very

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fast dissolution of gelatin, a mass of fibroblasts, monocytes, and multinucleated giant cells, as well as blood vessels, can easily grow into the inside of the membrane, resulting in a fast degradation speed for PG30 and PGH30. The HAC-mediated homogeneous composite is also helping for maintaining the membrane barrier function. The GTR membrane should preserve its shield function for 3–6 months to protect hard tissue regeneration. Both P0 and P30 maintained their block function for as long as 24 weeks. The slow degradation rate of P0 and P30 membrane suggested that the requirements of physical integrity during tissue regeneration were satisfied. However, gelatin incorporation dramatically 34

ACCEPTED MANUSCRIPT increased degradation. According to our results, although the degradation rate of the HAC-mediated homogeneous membrane was slower than that without the HAC sample, while the gelatin reached 50 wt% of the total weight of gelatin and PCL, the membrane barrier function can only be maintained for

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8 weeks. Thus, gelatin content must be reduced for GTR/GBR application. 5. Conclusions:

This study comprehensively and systematically investigated the structure and performances of 4

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types of PCL-based electrospun nanofiber membranes to select a suitable material system for

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application in GTR/GBR treatment. The TS of all the membranes satisfied the requirement of the GTR membrane. The PCL-based polymer matrix was suitable for preparing the MNA loading GTR/GBR membrane because the drug encapsulation efficiency was relatively high and the drug release profile was consistent with the inflammation cycle characteristics. Incorporation of gelatin and MNA greatly

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improved hydrophilicity, leading to improvements in cell biocompatibility. Trace amounts of HAC effectively prevented phase separation in spinning solution, resulting in homogeneous gelatin/PCL nanofibers and uniform drug dispersion. Drug encapsulation efficiency increased and the drug release

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speed decreased with HAC incorporation. A high local drug concentration and the very small amount

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of HAC caused no adverse tissue reactions. Thus, the PCL/gelatin/MNA/HAC composite is very suitable for preparing the GTR/GBR membrane. However, the proportion of gelatin should be reduced to maintain membrane integrity within 3 months. Membrane performance should be further confirmed in a periodontal regeneration infection model. Additionally, this drug delivery system may be broadly applicable to any therapeutic approach in which controlled drug delivery is necessary, such as wound healing, prevention of post-surgical adhesions, and tissue engineering applications.

Acknowledgements 35

ACCEPTED MANUSCRIPT This work was supported by the National Natural Science Foundation of China (51303014, 81330043), Beijing Nova Program (Z131102000413015), National Outstanding Youth Science Fund (50725310), The Ministry of Science and Technology of the China (2012BAI10B02), and Beijing

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Municipal Training Programme Foundation for the Talents (2013D00303400041).

Supplementary data

Supplementary data associated with this article can be found, in the online version, at …..

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[30] Zeng J, Yang LX, Liang QZ, Zhang XF, Guan HL, Xu XL, Chen XS, Jing XB. Influence of the drug compatibility with

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polymer solution on the release kinetics of electrospun fiber formulation. J Control Release, 2005; 105 : 43–51. [31] Fang ZD, Fu WG, Dong ZH, Zhang XM, Gao B, Guo DQ, He HB, Wang YQ. Preparation and biocompatibility of electrospun poly(l-lactide-co- -caprolactone)/fibrinogen blended nanofibrous scaffolds. Appl Surf Sci, 2011; 257: 4133–4138. [32] Wu L, Li H, Li S, Li XR, Yuan XY, Li XL, Zhang Y. Composite fibrous membranes of PLGA and chitosan prepared by coelectrospinning and coaxial electrospinning. Biomed Mater Res, A, 2010;92: 563–574. [33] Zhou S, Peng H, Yu X, Zheng X, Cui W, Zhang Z, et al. Preparation and characterization of a novel electrospun spider silk fibroin/poly(d, l-lactide) composite fiber. Phys. Chem. B，2008; 112: 11209–16 [34] Xiang A, McHugh AJ. A generalized diffusion–dissolution model for drug release from rigid polymer membrane matrices, J Membr Sci 2010; 366: 104–115. [35] Xiang A., McHugh AJ. Quantifying sustained release kinetics from a polymer matrix including burst effects, J. Membr. Sci. J. 2011;371: 211-218. [36] Kim K, Luu YK, Chang C, Fang DF, Hsiao BS, Chu B, Hadjiargyrou M, Incorporation and controlled release of a hydrophilic antibiotic using poly (lactide-co-glycolide)-based electrospun nanofibrous scaffolds, J Control Release 2004;98:47 –

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poly(lactide-co-glycolide) microspheres in vivo and in vitro, Biomaterials , 1999;20:1057-1062

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Table

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Tables Table 1 Electrical Conductivity of the Spinning Solution as well as the Pore size and porosity of membranes Pore size （μm）

Porosity (%)

Electrical Conductivity(μs/cm)

P0 P30 PG30 PGH30

6.04±0.64 3.04±0.50 3.21±0.80 4.01±0.70

76.4 70.2 70.5 70.9

6.43±1.32 7.41±1.48 18.63±2.25 19.04±1.39

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Sample

EB(%)

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Table 2. TS and EB of the different PCL-based membranes in dry and wet state

TS (Mpa)

EB(%)

325±35

12.55±2.82

152±16

99±27

11.84±1.27

52±16

19±3

6.23±1.53

13±4

24±6

3.88±0.49

30±9

TS (Mpa) P0

9.99±1.02

P30

6.33±0.84

PG30

4.14±0.90

PGH30

4.26±1.17

Wet State

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Dry State

Samples

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Table 3 Water contact angle values of electrospun PCL-MNA nanofiber membranes Samples

CA(°)

P0

129.64

P30

83.99

PG30

36.69

PGH30

61.16

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Supplementary material

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Supplementary Information S1 materials PCL pellets (Mw = 80,000) and 3-(4, 5-dimethylthiazol-2-yl-2, 5-diphenyltetrazolium bromide) (MTT) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Gelatin with biomedical grade was

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obtained from Rousselot (France). Metronidazole (MNA) was purchased from Tokyo Chemical

Industry Co., Ltd. (Japan). Dulbecco’s modified eagle medium (DMEM), Fetal bovine serum, 0.05%

Trypsin EDTA and phosphate buffer saline (PBS) were purchased from Gibco (USA). Methanol with HPLC reagent was purchased from Alfa-Aesar Chemical Inc. (USA). Other chemical reagent were

purchased from Sinopharm Chemical Reagent Beijing Co., Ltd. (China) and used as received without

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further purification. L929 fibroblast cell lines were kindly donated by Jishuitan Hospital.

S2 Composition and preparation of the spinning solution

Spinning liquid of P0: The solution of PCL with concentration of 10 wt% was prepared by dissolving PCL in DCM: DMF (60:40 v/v) mixtures and stirring magnetically for 24 h at room

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temperature. Spinning liquid of P30：The PCL solution with concentration of 10 wt% was prepared by dissolving PCL in DCM: DMF (60:40 v/v) mixtures and stirring magnetically for 24 h at room temperature. MNA, which was dissolved in DCM with the concentration of 30% w/w with respect to the polymer used, was added to the polymer solution and stirred together for another 12 h to get a continuously electrospinning solution. Spinning liquid of PG30：PCL/HFIP (6 wt.%) and gelatin/HFIP (6 wt.%) solutions were separately prepared under magnetic stirring for 12 hours at room temperature. Both solutions were then mixed to obtain the solution of PCL/gelatin at the mass ratio of 50:50. MNA, with concentration of 30% w/w with respect to the polymer used, was added to the polymer solution

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and stirred together for another 12 h to get a continuously electrospinning solution. Spinning liquid of PGH30：HAC, with concentration of 0.1% v/v with respect to the total volume of polymer solusion gotten from electrospinning solution of PG30, was added to the polymer solution and stirred together for another 12 h to get a continuously electrospinning solution. The solution was electrospun from a 20ml syringe with a needle diameter of 0.4 mm and fed to the needle tip at a rate of 1 ml/h by a syringe pump. Electropinning voltage was applied between the

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needle and the grounded collector which was laid with aluminum foil at a rotating rate of 300 rpm. The needle was located at a distance of 20 cm from the ground collector. Optimized high voltage (8-12 KV) was applied to the solution with different content of drug to get a stable polymer jet. All experiments

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were carried out at room temperature and relative humidity of 40%.

S3 Equations for Scaffold apparent density and Scaffold porosity Scaffold apparent density (g/cm3) =

Scaffold porosity (%) = (1 


   

      

   !⁄"#$ & '(  /"#$

 100.

, . S1

(eq.S2)

S4 Molecular Dynamics simulation Approach and the interaction between PCL polymer matrix and MNA In this force-field approach, the total energy ET of the system is represented by the sum of

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bonding and nonbinding interactions given as ET=Eb+E0+Eψ+Eloop+Epe+Evdw+Eq. The first five terms represent the bonding interactions, which correspond to energies associated with the bond Eb, bond angle bending E0, torsion angle rotation Eψ, out of loop Eloop, and potential energy Epe. The last two terms represent the nonbonding interactions, which consist of the van der Waals term Evdw and electrostatic force Eq. This force-field approach describes the intramolecular and intermolecular interactions in the electrospinning system of MNA and PCL. The COMPASS force field

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has been widely used to optimize and predict the structural, conformational, and thermophysical condensed phase properties of molecules including polymers. Initial velocities were set by using the Maxwell-Boltzmann profiles at 298 K. The Verlet velocity time integration method was used with the time step of 1 fs.

Amorphous cells were constructed containing blends of four chains of 30 repeat units of PCL and

MNA with different content. First, the PCL polymer chains and MNA were built in periodic boundary

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cell (Fig.S1). Then the cell was taken by energy minimization to get the cell with the lowest energy. After that, the systems were equilibrated in the isothermal-isobaric (NPT) ensemble at 298 K. This equilibration was usually done for 5 ps with the dynamics that was followed by the data accumulation

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to analysis hydrogen bonds.

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running for at least 100 ps with the configurations saved every 5 ps. At last, the proper cell can be used

Fig. S1. Simulation for hydrogen bonding through molecule modeling using Materials Studio Software

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(in the unit cell: red, O; white, H; gray, C; blue, N; purple, back bonds of PCL matrix).

S5 FTIR, DSC and XRD analysis

FTIR analysis was performed on a Bruker Tensor 27 FTIR spectrometer. The scan range was 4000

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cm−1 to 600 cm−1 with a resolution of 2 cm−1. The thermal properties of the membranes were investigated by the DSC measurements (Mettler-Toledo International Inc., Switzerland). Samples with a weight of approximately 5 mg were loaded in an aluminum crucible under dry condition. The samples were cooled from 30 oC to -100 oC and stayed for 5 minutes, then heated to 220 oC under nitrogen atmosphere at a heating rate of 10 oC /min. XRD analysis was carried out on a Rigaku D/max-Ultima III X-ray diffractometer equipped with Cu-K α source and operating at 40 kV and 40 mA. The diffraction patterns were obtained at a scan rate of 5 °/min. The mechanical properties of the membranes both at dry state and wet state were also evaluated .by using a BOSE ElectroForce 3200 test instrument with a 50-N load cell at a crosshead speed of 5 mm/min at the ambient temperature of 25℃. All samples were cut into rectangles with dimensions of 25 mm×4 mm. Five samples at dry state were tested for each membrane. Another five samples were soaked in deionized water for 1 h for wet state tensile property measurements. The thicknesses of the samples were measured with a

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micrometer accurate to 1 µm.

S6 Cytotoxicity evaluation MTT assay was used to test the cytotoxicity of the membranes with different content of MNA. The 20 mm30 mm samples were sterilized by washing with a 75% (v/v) ethanol solution in sterilized water, exposed to Co60 for 15 min, and then incubated in DMEM at a proportion of 3 cm2/mL for 24 h at 37 oC to get the extract solution of the samples. The extract solution was then filtered (0.22 µm pore

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size) to eliminate the possible presence of solid particles of the sample. L929 cells were resuspended in DMEM supplemented with 10% (v/v) fetal bovine serum (FBS) at a density of 1.0 104 cells/mL and

100 µL cell resuspension solution was pipetted into 96-well micrometer plates. After incubated at 37 ℃ under 5% CO2 atmosphere for 24 h, the medium was replaced by the previously prepared extracted dilutions (50%, volume proportion), the culture medium as blank control and DMSO as negative

control. After 24 h, 48 h and 72 h of incubation, morphology of the cells in plate were observed using

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inverted phase contrast microscope (Olympus IX50-S8F2), then the cells were treated with 20 µL/well MTT (5 mg/mL in PBS solution) and incubated for another 4 h at 37 oC in a humidified atmosphere of 5 % of CO2. At this stage the MTT and culture medium were removed and 200 µL/well of DMSO was

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added to dissolve the forming formazan crystals. After shaking the plate for 15 minutes, the optical density (O.D.) was read on a multi-well microplate reader at 630 nm. All materials’ extracts were tested for six averages.

Cytotoxicity of P30 was also tested with human periodontal ligament fibroblasts (hPDLFs) and ROS cells using a Presto Blue assay (Life Technologies). PrestoblueTM is a blue non-fluorescent, cell permeable compound (resazurin-based solution), that is reduced by living cells into a fluorescent compound (resorufin). Commercially available hPDLFs were cultured in DMEM supplemented with 10% fetal calf serum (FCS), 2 mM glutamine, 50 U/mL penicillin and 50 U/mL streptomycin. ROS

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cells were cultured in α-MEM supplemented with 10% (v/v) FBS. P30 was cut into 20 mm30 mm rectangle. The samples were disinfected in 100% isopropanol for 20 minutes under shaking followed by three rinses in PBS under sterilization environment. After that, 8 mL cells culture medium (DMEM+10 % FCS for hPDLFs and α-MEM+10 % FBS for ROS cells) were separately added into the sample in a sterilization tube to get the extract substrates at 37 oC for 72 hours. Cells were resuspended accordingly in culture medium at a density of 5105 cells/mL and 100 µL cell resuspension solution

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was pipetted into 24-well plate after 1 mL culture medium was added in the plate. After incubation for 24 hours, medium were replaced with 900 µL extract substrates per well. The viability of the cells cultured in the extract solution was assessed using the PrestoBlue assay. With another 24 hours

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incubation, 100 µL PrestoBlueTM was pipetted per well. At 1 hour incubation, 200 µL dye-containing medium of each well was placed into 96-well plate in duplicate, and fluorescence intensity was measured using a fluorescence reader (TECAN Spectrophometer) at an excitation wavelength of 570 nm and emission wavelength of 600 nm. Every sample experimented for 3 averages.

S7 Proliferation of L929 cells on the nanofiber membranes The membrane was cut into circle 2.5 cm in diameter, sterilized, fixed on a 24-well plate Cell-Crown

(Sigma, USA), and then put into the well of 24-well plate. L929 cells were cultured in DMEM supplemented with 10% FBS and antibiotics (100 µg/mL penicillin and 100 mg/mL streptomycin) at a density of 4.0104 cells/mL. Cell culture medium(800µL) and cell suspension (100µL) were plated onto the sample carefully. Cell suspensions added into wells with no materials were regarded as control. Plates were incubated at 37 oC under 5% CO2 atmosphere. The medium was changed every two days. After incubating for 4 h, the sample was washed with PBS for three times to remove the cells didn’t

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attach on the membrane and 100 µL L WST-8 WST 8 (final dilution: 1:10) which can react with dehydrogenase in mitochondria getting a water soluble oluble formazan was added onto the sample to test the alive cells. After incubating for another 4 h, 100 µL of supernatants was transferred into a 96-well 96 well plate to read the optical density (O.D.)) at 450 nm. At days 1, 3, 5 and 7 of culture, cell proliferation proliferation and viability were also assessed using CCK-88 test as above. The cells-cultured cells cultured membrane was afterwards washed with PBS for three times, fixed with 3% glutaraldehyde at 4 oC for 2 h, soaked in 0.18 M sucrose solutions

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at 4 oC for 2 h and then was dehydrated ated through a series of graded ethanol solutions and lyophilized overnight. The morphology of cells on the membranes were observed by using SEM.

S8 Antibacterial activity

A 100 µL L aliquot of Fusobacterium nucleatum reconstituted in Brian Heart Infusion culture medium and previously subcultured was spread onto an agar plate. Sections (1.0 cm1.0 cm 1.0 cm) of

nanofiber membranes with different contents of drug were placed on agar plates (three averages for one

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sample on the same petri dish). The plates were incubated for different times at 37℃ 37 under anerobic conditons. The bacterial growth on the plate was visualized directly and the diameter of the inhibition

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zone was measured on days 1, 4, 7, 14, 21, and 30.

S9 histological study and SEM studies

For histological study, the membranes with surrounding subcutaneous tissues were removed. The specimens were fixed with 4% formaldehyde in PBS. After dehydration in a graded series of alcohol, the specimens were embedded in paraffin paraf wax and cut into 5µm m transverse sections. These sections were stained with hematoxyline-eosin eosin (H&E) for the observation by light microscopy (Model BX50F4, Olympus, Japan).

Explanted membranes and surrounding tissue used for SEM studies were washed three times in Dulbecco’s phosphate-buffered buffered saline (D-PBS; (D PBS without Ca2+ or Mg2+, pH=7.2) and fixed in 3%

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glutaraldehyde overnight prior to dehydration in increasing concentrations of ethanol (30%, 50%, 70%, 90%, 95% and 100%, 1h each). Then samples were left to air-dry air y in the fume hood for 1 h. Completely dried samples were sputter-coated coated with gold and observed by SEM.

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S10 Photograph of the PCL/GT and PCL/GT/HAC solutions

Fig.S2 Photograph of the PCL/GT after being left to stay 8h and PCL/GT/HAC solutions with a mass ratio of PCL / GT at 50:50 after being left to stay at rest for 24 h

S11 Formation of HBs between MNA and PCL matrix

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Fig.S3 Formation of HBs between MNA and PCL matrix of nanofiber membranes with different MNA content nt simulated by molecular modeling using Material Studio Software: (a) 1 wt.%; (b) 5 wt.%; (c) 20 wt.%; (d) 40 wt.%.

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S12 results of ATR-FTIR spectra, DSC measurement, measurement XRD investigation According to the ATR-FTIR FTIR spectra of the electrospun membranes (Fig.5a), (Fig.5a), the infrared spectrum of P0 is consistent with the PCL-related related stretching modes, including 2943 cm-1 (asymmetric CH2

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stretching), 2866 cm-1 (symmetric CH2 stretching), 1721cm-1 (C= O stretching), 1294 cm-1(C–Oand Oand C–C stretching) and 1240 cm-1(asymmetric C-O-C C stretching) bands. The MNA-related related stretching modes are presented by the peak at 3414 cm-1 and 3220 cm-1 (O-H stretching), 3097 cm-1 (=CH) and 1550cm-1 (-NO2 antisymmetric stretching) . The characteristic peaks of MNA emerged on the spect spectra of P30, PG30 and PGH30 are no different from those of pure MNA crystals crystals，demonstrating that the electrospinning process did not adversely affect the molecular structure of MNA，especially MNA the antibacterial -NO2 group. After the incorporation of MNA , the characteristic of C=O on PCL red shift from 1725cm-1 to 1721cm-1 , revealing the HBs between PCL and MNA formation, and consistent with the results of molecular dynamics simulation . Common bands of protein related peaks at approximately 1650cm-1 (amide I) and 1540(amide Ⅱ),corresponding ),corresponding to the stretching vibrations of C=O bond, and coupling of bending of N-H N bond and stretching of C-N bonds respectively, are appeared on the spectra of PG30 and PGH30. Two strong absorption bands at 1716 cm−1 and 1273 cm−1

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appeared in the IR spectrum of PG30, corresponding to the characteristic bands of PCL originally situated at 1725 cm−1 and 1294 cm−1, and showed a measurable shift towards relatively lower wave numbers. Besides this, the characteristic amide bands of gelatin at 1650 cm−1 and 1540 cm−1 showed slightly sharp intensity and shifted to around 1643 cm−1 and 1538 cm−1 respectively. The shifting of original absorption bands towards the lower wave numbers in PG30 and PGH30 indicates that HBs exist between PCL and gelatin moleculers. FTIR spectrum of PGH30 was nearly same as that of PG30.

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The consistency of the infrared spectrogram also demonstrated that the acetic acid didn’t react with

other components during electrospinning process. That is because the surface of PG30 used for FTIR test was fabricated in the early stage of electrospinning, and the electrospinning solution was

heterogeneous at that moment, so the material composition of PG30 was just as same as PGH30, resulted in the same FTIR spectrum pattern.

DSC measurement was used to investigate the crystallinity of PCL, drug distribution and acetic

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acid-mediated homogeneity of the nanofiber membranes. DSC thermograms of MNA, P0, P30, PG30 and PGH30 are shown in Fig. 5b. The crystalline melting peak of pure PCL occurs at 60℃. With addition of MNA (P30), a second melting point was detected at 144℃, which is lower than the melting

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point of pure MNA at 162℃(Table S1)，indicating a part of the MNA molecules interacted with the PCL. The melting enthalpy (∆Hm) and the crystallinity of PCL matrix decreased after the MNA was incorporated. When gelatin was also added, the ∆Hm and the crystallinity of PCL matrix further reduced compared to those of P30. However, ∆Hm and the crystallinity of MNA increased and melting point had a little high shift. As the case of PGH30, ∆Hm and the crystallinity of PCL matrix decreased, while the ∆Hm of MNA also decreased when compared to PG30.

Table S1 melting enthalpy and crystallinity of PCL-based nanofibermembranes

P0 P30 PG30 PGH30

∆Hm (J/g)

Xc (%)

89.46

68.8

55.70

40.1

46.45

35.7

23.77

17.1

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XRD are also applied to investigate the microstructure of the membranes. In the XRD pattern of P0, a sharp peak at 21.3° and a relatively low intensity peak at 23.6°are assigned to the characteristic peaks of semi-crystalline PCL. When 30% percent of MNA was added (P30), the crystal peaks of MNA located at 12.2° and 13.8° appeared which demonstrated the aggregation of MNA. As the incorporation

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of MNA, the intensity of characteristic peaks of PCL decreased indicates the reduction in crystallinity of PCL. Besides, the peaks shift to higher angle slightly, indicating the length of the PCL crystal stacks decrease, probably due to the MNA molecules among the PCL molecular chains restrain the formation of PCL crystal stacks and the orientation of PCL molecular chains also affects the PCL crystallization. . Pure gelatin showed no peak in XRD pattern, which indicates its amorphous nature. The addition of gelatin resulted in the lower intensity of the characteristic peak of PCL compared to P30. However, the intensity of the characteristic peaks of MNA at 16.2°，17.2°, 18.2°, 25.2°, 27.4°, 28.1°, 29.4°, and

29.9°increased.

Higher of the crystal peak value implied that more MNA aggregated in PG30

compared to P30. The X-ray diffraction of PGH30 is very similar with that of PG30. The difference is that the intensity of the characteristic peaks of MNA in PGH30 is lower than those in PG30, which indicated less MNA aggregation in acetic acid adding sample than sample without acetic acid.

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S13 Surface hydrophilicity

Fig.S4 Water drop on the surface of the membranes

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P0 has the highest CA, with the 30wt% MNA addition, the CA value of the membranes showed a significant decrease from 129.64°to to 83.99°, 83.99 indicating the obvious improvement of the hydrophilicity. hydrophilicity The CA of the membrane also greatly decreased from 8 83.99° to 36.69° with the incorporation of gelatin

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suggesting the he hydrophilicity of the membranes greatly improved. Results from water contact angle measurement confirmed that both of the MNA and gelatin can improve the hydrophilicity of the nanofiber membrane significantly the CA value showed an increase manner compared to PG30, suggesting a little decrease of the hydrophilicity.

S14 In vitro biodegradation

For P30, the drug encapsulation efficiency was 83.2%. according to drugs accounted for 30% of the weight of PCL , The he weight percentage of drug /membrane was calculated as 30×83.2%/ 83.2%/ （30+30/30%）=19.2%, =19.2%, so the mass loss of membrane resulted by the release of drug was 19.2%. The

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weight loss data from weighing method showed that the weight loss was about 19 19.1% .1% at the 14th day. Because the PCL do not degaradate in the first two weeks, so the degradation is mainly caused by the drug release.

For PG30, The drug encapsulation efficiency of PG30 was 84.9%. The weight percentage of drug to membrane was calculated as 30×84.9%/ 84.9%/（30+30/30%）= = 19.6%. The weight percentage of gelatin was calculated as (30×0.5/30%)/（30+30/30%） 30+30/30% =38.5%. The total weight percentage of the gelatin and

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MNA was calculated as 19.6%+38.5%=59.1%.

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Fig.S5 surface morphology of PCL-based PCL membranes after 30 days degradation

S15 Mechanical properties of the nanofiber membranes after degradation for 1 month

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Table S2 construction of the mechanical properties of different membranes after before and after degradation in PBS solution for 30 days. Sample

EB /%

before

after

before

after

9.99

6.34

325

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P0

TS /MPa

6.33

5.57

99

50

PG30

4.14

3.43

19

21

PGH30

4.26

2.49

24

8

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P30

S16 cell relative growth rate (RGR) RGR reflects the viability of L929 cells growing in the material extracts. The TCP was selected as

the control material. The RGR was classified to six grades according to the standard GB/T16175 GB/T16175-1996 1996. The RGR has been divided into six classes as ≧100%, 75~99%, 50~74%, 25~49%, 1~24% and 0,, corresponding to Grades 0 to 5. Grades 0 and 1 were accepted as qualified. Grade 2 should be considered by combining with the cell’s morphology. Greater than grade 3 was considered as unqualified.

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Absorbance of the materialmaterial Absorbance of the blank RGR (%)= Absorbance of the negativenegative Absorbance of the blank

(eq.S3 3)

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Cell relative growth rate and the cell toxicity grades are classified to grades according to the

standard GB/T16175-1996 1996 (see S11). After 24 hours of incubation, the RGR of P0, P30 and PG30

reached to 100% or even higher than 100%, should be classified to grade 0. There are no significant

differences between blank and experimental groups from statistics calculation, which reflects the good cytocompatibility of all nanofibers. RGR of PGH30 was 84%, which is little smaller than that of other

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samples, but stilll achieved the level of grade 1.

Fig. S6 Microscope images of the L929 cells cultured in the extract substrates of PCL-based

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electrospun membranes after 72h

S17 cytotoxicity of P30 with hPDLFs and ROS cells

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membrane for 24 hours.

S18 SEM studies detected on days 1, 3, 5 and 7

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Fig. S7. Fluorescence of hPDLFs cells and ROS cells cultured in extract substrates of P30 electrospun

To observe the effects of different components on of fibroblasts, SEM studies were carried on

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membranes were detected on days 1, 3, 5 and 7. The cellular morphology on different membranes on the 1st and 7th day is shown in Fig.10.c and d. On the 1st day, most of the cells presented typically round shape on P0 and P30, but spindle shape on PG30 and PGH30. The specific structures related to the cell motility of fibroblasts— filopodia and lamellipodia—can be observed. It seems like that more interactions and contacts among cells on PG30 and PGH30 than those on P0 and P30 on the first day. On the 7th day, the cells observed to be characteristically spindle shaped and stretched across the nanofibrous substrates on all membranes. For gelatin containing samples, the cells form almost a confluent layer on the nanofiber membranes. The cells emitted cytoplasmic process towards the

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nanofibers and neibouboring cells, forming communication with the surrounding micro-environment and neighboring cells. It was estimated that only 50%-60% L929 cells confluence with most round shaped cells on P0 and P30, but reaching an estimated confluency of about 90% on PG30 and PGH30 on the 7th day. It seems like that the acetic acid in PGH30 also had no side effects on cell proliferations.

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S19 Drug release calculation Clinically, the recommended oral dosage of MNA for an adult patient is 0.4 g-0.8 g every time which is much larger than the minimal inhibitory concentration value (0.2-2 mg/L) used to inhibit the growth of aerobic or anaerobic microorganisms. According to the density of the electrospun membranes (~0.24 g/cm3) and the prototypical mediated dimension of the GTR membrane (5cm×5cm×250µm), for P30, we can get the total amount of MNA is 0.288 g. It can be deduced that the concentration of drug released around the membrane is high enough to inhibit the microorganisms growth and much lower than the systemic treatment dosage.

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